1. Field of the Invention
The present invention relates generally to an imaging device, and more particularly to a scintillation crystal for an imaging device.
2. Description of the Background Art
Radiographic imaging is the detection of radiation in order to form an image. By detecting the amount of radiation passing through or emanating from a test subject, the resultant image may give a representative view of the materials and construction of the test subject.
Gamma rays are a form of radiation that is emitted by excited atomic nuclei during the process of passing to a lower excitation state. Gamma radiation is capable of passing through soft tissue and bone. Gamma radiation may therefore be used for medical imaging, among other applications.
Gamma radiation for medical imaging usually involves a radiopharmaceutical, such as thallium or technetium, for example, that is administered to the patient. The radiopharmaceutical travels through the patient's body, and may be chosen to be absorbed or retained by an organ of interest. The radiopharmaceutical generates a predictable emission of gamma rays through the patient's body that can be detected and used to create an image. This may include imaging areas of the body or imaging specific organs. For example, a radiographic imaging device may be used to capture images of the heart, including real time images.
A radiographic imaging device may be used to detect radiation emanating from the patient and may be used to form an image or images for viewing and diagnosis. The radiographic imaging device may be a device such as a gamma or gamma ray camera, also referred to as a scintillation camera or an Anger camera. The radiographic imaging device allows a doctor to perform a diagnosis on a patient in a non-invasive manner and additionally may allow the doctor to observe organ function. In addition, the radiographic imaging device may be used for other imaging functions.
FIG. 1 shows a typical gamma camera or gamma camera component unit, including a scintillation crystal 105 having an emission face 108, a sensor 112 or a sensor array, and a processing apparatus 116.
The scintillation crystal 105 is typically a thallium doped sodium iodide crystal (NaI(Tl)) that generates photons upon absorbing radiation particles, such as gamma radiation (i.e., it scintillates). This scintillation process converts the gamma radiation into light photons, which can be more easily detected. The photons emerge from the emission face 108 and may be detected by the sensor 112 or sensor array.
The sensor 112 may be, for example, a photon detection sensor such as a photomultiplier tube. The sensor 112 receives the photons generated by the scintillation crystal 105 and converts them into a representative electronic signal. A typical photomultiplier tube 112 may include a semi-transparent photocathode, a focusing grid, dynodes, and an anode (not shown). Multiple sensors 112 may be used to form a sensor array in a radiographic imaging device.
The processing apparatus 116 receives the electronic signals from the sensor 112 or sensor array, may amplify and filter the signals, and processes them to form an image.
FIG. 2 is a block diagram of a scintillation crystal 105 and an associated array of sensors 112, as used in a gamma camera, for example. Each sensor 112 generates an output current signal that may be amplified, filtered and processed to generate an image. The scintillation crystal 105 emits light photons upon absorption of radiation, thereby generating an event 122. For example, the event 122 may emit a plurality of photons that may be received by one or more of the sensors 112. It should be understood that the event 122 may be any type of event, including a source of photons, a source of rays due to a radioactive decay, a source of electrons or protons, a source of electromagnetic waves, etc.
FIG. 3A shows a histogram created by photon counts from the sensors 112 in a direction along the X-axis in response an event 122. A histogram may be thought of as a visual representation of the contents of a series of storage bins that count photons as they are received and categorize them according to the location of the sensors along a predefined direction. The histogram therefore creates a wave form peak. The peak may indicate the approximate center of the event 122 (e.g., a centroid). The centroid may be determined from one or more sensor outputs and may be used to create an image composed of one or more event centroids. Similar considerations apply to the Y-centroid, as shown in FIG. 3B.
FIG. 4 shows an array of sensors 112 and the event 122. Due to the difference in spacing and distribution along the X axis and along the Y axis, as is reflected in the distances DX and DY, it is highly desirable that some correction or adjustment be made to the sensor readings in order to accurately determine the centroid of the event 122.
The figure shows an example of a hexagonal sensor array used for imaging a large area. Typically, rows of the contiguous sensors 112 are aligned in the direction of a Cartesian axis of the scintillation crystal, such as the X-axis, for example. However, if the sensors 112 are grouped into an array in a square grid fashion, large gaps are left between the individual sensors 112, as the sensors 112 are typically round. Therefore, in order to minimize the gaps and pack the sensors 112 together as densely as possible, the sensors 112 are typically formed into a grid having a substantially hexagonal arrangement (as shown) in order to minimize gaps between the sensors 112.
FIG. 5 shows the event 122 and the sensors 112 that are affected by the event 122. From FIG. 5 it can be seen that in the X dimension, there are approximately five samples (columns of sensors) within the area of the event 122. However, in the Y dimension, there are only approximately three samples (rows of sensors) within the area of the event 122. Because the sensors 112 are hexagonally packed as shown in FIG. 4, the histogram in the X-axis direction will have more bins than the Y-axis histogram (see FIG. 3B), and therefore resolution and linearity along the Y-axis will suffer in comparison to the resolution and linearity along the X-axis.
In scintillation crystals of approximately 2.5 cm thickness, the difference in sampling along the X- and Y-axes produces a measurable and problematic difference in resolution and linearity characteristics in the X and Y directions. This effect is due to the differences of sensor spacing along the X-axis and the Y-axis, as can be seen from the markings along each axis as shown in FIG. 5. The spacing of the X-axis samples is 1*R, where R is the photomultiplier tube radius; while the Y sample spacing is (√{square root over ( )}3*R) or (1.73*R). Therefore, when using standard positioning algorithms for generating an image, different resolution and linearity characteristics are obtained in the X and Y directions.
Early gamma cameras utilized scintillation crystals approximately 1 cm in thickness. For such crystals, the difference in X-axis and Y-axis sampling produces only minor differences in resolution and linearity characteristics. Recently, however, the crystal thickness has been increased to more than 2.5 cm in order to facilitate detection of high energy radiation. The extra crystal thickness is needed because the high energy radiation penetrates farther into the crystal, while the lower energy radiation does not penetrate very far. The lower energy radiation therefore causes scintillations that are farther from the emission face 108, and the resulting photons may spread more within the scintillation crystal 105 before reaching the emission face 108 (see FIG. 2).
A prior art approach to radiographic image improvement has employed a series of channels cut or otherwise formed in the scintillation crystal. The purpose of the channels is to guide emitted photons through the crystal and to minimize the lateral movement (“spreading”) of photons within the crystal. By minimizing spreading, the sensors are more likely to receive photons emanating from scintillation events located directly below them. The channels therefore function somewhat as a collimator and direct or channel the photons in a direction substantially perpendicular to the plane of the scintillation crystal 105 and/or perpendicular to the emission face 108.
In one prior art approach, the scintillation crystal 105 is formed with two sets of substantially orthogonal channels which form a plurality of rectangular solid portions. The channels of the prior art scintillation crystal 105 are formed to be of identical and uniform depths. The prior art approach is used by Bicron, Newbury, Ohio, 44065, in their STARBRITE crystal.
However, for a channel depth sufficient to achieve good performance along the X-axis (i.e., to achieve an acceptable output resolution and an acceptable output differential linearity), the corresponding resolution and differential linearity along the Y-axis is considerably worse. The degradation in performance is attributable to the coarser sampling of the light response function, especially in the tail of the light response function along the Y-axis. The fall off of the light response function with distance is quite sharp because of the depth of the grooves (approximately one-half of the crystal thickness).
There remains a need, therefore, for improvements in radiographic imaging devices.